This invention relates to coherent imaging methods applicable to a phased array ultrasonic transducer and, more particularly, to beamforming techniques for use in ultrasound imaging systems.
X-ray mammography is presently the primary screening procedure for detection of breast lesions. Sonography is often used to confirm screening results and also used as a complementary method when X-ray mammography has failed to confirm the results of a manual examination. Often, two separate imaging procedures are required for a single patient, which is inconvenient and may even delay diagnosis.
Hand-held ultrasound transducer probes have been used in free-hand examinations to complement X-ray mammography. A drawback of such freehand examinations, when used to supplement mammography, is the inability to provide geometric registration between the mammogram and ultrasound images. This lack of registration makes it difficult to relate what is seen in the ultrasound image to what is seen in the mammogram. Furthermore, the three dimensional shape of the lesions and the increased vascularity associated with carcinoma make volumetric spatial registration of the ultrasonic data with a mammogram desirable. In light of the foregoing, the development of equipment capable of performing both imaging exams, i.e., X-ray and ultrasound, at the same time and in registration would be an advance.
U.S. Pat. No. 5,479,927 describes apparatus that combines mammography equipment with an ultrasonic transducer to generate ultrasonic images of the internal structure of breast tissue that are in geometric registration with a mammogram. The apparatus includes a radiolucent and sonolucent compression plate. Either before or after the X-ray exposure, a carriage-mounted ultrasound transducer is translated in increments across the compression plate to generate a plurality of sectional views of the breast tissue. The X-ray and ultrasound images produced by this sonomammography apparatus are ideally in geometric registration. Those images may in turn be processed by a conventional microprocessor-based workstation to provide holographic views of the internal features of a patient""s breast.
X-ray mammography images are typically obtained using a plastic plate to compress the breast. The compression plates used in X-ray mammography were historically made of polycarbonates because of their tensile strength and transparency to X-ray. These materials are acoustically opaque. Because of their high densities, most other materials potentially useful for the compression plates in mammography equipment have relatively high attenuation and reflection coefficients for acoustic wave energy. U.S. Pat. No. 5,479,927 describes use of a compression plate made of a high-performance acoustically transparent (sonolucent) and X-ray transparent (radiolucent) film which is sufficiently rigid to serve as a compression plate at a thickness of about 25 micron (1 mil).
Since it would be advantageous to have images that are directly comparable for the X-ray and ultrasound modalities, it would be desirable to form the ultrasound image through the compression plate. However, since acoustic propagation within the compression plate is substantially different than in water or the coupling gel, refraction effects on the waves emitted from the transducer elements of a phased array would severely corrupt the beamforming process of an ultrasound imager, since the beamforming time delays assume a constant velocity of 1,540 m/sec. (meters per second).
Conventional ultrasound imaging systems comprise an array of ultrasonic transducer elements which are used to transmit an ultrasound beam and receive the reflected beam from the object being studied. Such scanning comprises a series of measurements in which the focused ultrasonic wave is transmitted, the system switches to receive mode after a short time interval, and the reflected ultrasonic wave is received, beamformed and processed for display. Typically, transmission and reception are focused in the same direction during each measurement to acquire data from a series of points along an acoustic beam or scan line. The receiver is dynamically focused at a succession of ranges along the scan line as the reflected ultrasonic waves are received.
For ultrasound imaging, the array typically has a multiplicity of transducer elements arranged in one or more rows and driven with separate voltages. By selecting the time delay (or phase) and amplitude of the applied voltages, the individual transducer elements in a given row can be controlled to produce ultrasonic waves which combine to form a net ultrasonic wave that travels along a preferred beam vector direction and is focused at a selected point along the beam. The beamforming parameters of each of the firings may be varied to provide a change in maximum focus or otherwise change the content of the received data for each firing, e.g., by transmitting successive beams along the same scan line with the focal point of each beam being shifted relative to the focal point of the previous beam. For a steered array, by changing the time delays and amplitudes of the applied voltages, the beam with its focal point can be moved in a plane to scan the object. For a linear array, a focused beam directed normal to the array is scanned across the object by translating the aperture across the array from one firing to the next.
The same principles apply when the transducer probe is employed to receive the reflected sound in a receive mode. The voltages produced at the receiving transducer elements are summed so that the net signal is indicative of the ultrasound reflected from a single focal point in the object. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delay (and/or phase shifts) and gains to the signal from each receiving transducer element.
In a conventional ultrasound imager, the transmit and receive time delays are pre-computed and then stored in memory. These time delays are computed to compensate for time-of-flight from the center of the array element to the focal point, under the assumption of a uniform medium with a speed of sound of 1,540 m/sec. The coherence requirements for beamforming are quite demanding. Typically only about xc2xc cycle of error (peak to peak) can be tolerated before there is significant processing loss in signal and reduction in the resolution of the point spread function. For typical high-quality imaging the total error allowed is about {fraction (1/16)} of a cycle.
The time delays needed for coherent beamforming can be determined using the geometrical theory of diffraction in conjunction with ray tracing. The receiving array is attempting to dynamically focus on a moving wave packet that is traveling out from the array along a scan line anchored at the phase center with an angle xcex8 and at a constant speed. As the wave packet moves out, it illuminates reflecting elements that radiate a multiplicity of spherical waves back to the array. The array itself consists of discrete receiving elements that are arrayed in space according to the geometry of the transducer (e.g., linear arrays, curved arrays, two-dimensional arrays).
The basic nature of the diffraction involved in phased array beamforming can be reduced to a triangle. The hypotenuse of the triangle is determined by the path of a beam leaving the array at the phase center and traveling out at angle xcex8. The base of the triangle is the straight line distance between the phase center and the i-th transducer element in the array. The third side of the triangle is the distance from a point along the beam to the i-th element. As the wave packet moves out, its range R changes linearly with time according to the speed of propagation in the medium. At any instant, a reflection from a scatterer along the beam can radiate a spherical wave back to a receiving element. The triangle represents the fundamental geometry of a transducer. The distance from the phase center of the array to the field point (the location of the scatterer) is R and the location of the field point is easily given using sine and cosine of xcex8. To compute the distance from the field point to the i-th transducer element in the array, one must first calculate the values of the coordinates of the field point. The propagation distance is a function of the location of the field point and the location of the i-th element.
The foregoing triangulation method for computing beamformer time delays works well in the absence of an intervening layer having a speed of ultrasound propagation different from that of the biological tissue being imaged. However, since the speed of sound in a plastic mammography compression plate is typically much higher than the nominal speed of sound conventionally used to pre-compute the focusing delays of the ultrasound device, defocusing of the ultrasound beam will result. In order to produce a focused beam under these conditions, the transmit and receive time delays must be re-computed.
Passing a focused beam through a flat plate having different propagation speed from that of the surrounding medium need not cause noticeable defocusing. If the beam originates from a small aperture, or if the plate is thin (as is true case for an array with a thin plastic lens on its front face), then the effect of the plate is only to change the focal length by an amount related to the plate thickness. (In this instance, the flat plate can also be used to displace the focus from side to side, as is done in telescopes.) However, if the aperture is large (that is, if the maximum slope angle of the beam to the focus is large) or if the plate is thick, then spherical aberration results. Because of the short focal lengths and relatively large transducer apertures in an ultrasound imaging system, this problem will arise whenever a system operator attempts to acquire an ultrasound image of a compressed breast through a plastic compression plate. In that event, a suitable correction must be applied in order to form an in-focus image.
In a preferred embodiment of the invention, a method of imaging tissue comprises the steps of: compressing the tissue by pressing a plastic plate against the tissue; holding the compression plate in a fixed position; acquiring an X-ray image of the compressed tissue through the compression plate in the held position; acquiring an ultrasound image of the compressed tissue through the compression plate in the same held position; and registering the X-ray and ultrasound images.
The invention enables an ultrasound imaging system to acquire an image of tissue through a plastic plate. In a preferred embodiment, a beamformer is programmed with pre-stored transmit and receive time delays which have been computed to correct for the effects of refraction caused by an intervening plastic mammography compression plate of an X-ray mammography unit. This correction enables acquisition of an in-focus ultrasound image taken under the same conditions as an X-ray mammography image. Since the two images are formed from the same source under the same conditions, they can be registered, and their information compared on a point-by-point basis.
To eliminate the spherical aberration due to an intervening layer of plastic between the beamformer and the biological tissue being imaged, a ray tracing method is substituted for the conventional triangulation method in the computation of beamformer time delays. For the purpose of computing time delays, this approach uses a ray-tracing method where the refraction occurs at the interface of the plastic layer and adjacent tissue.